Differential phase contrast imaging constitutes an imaging method which in particular in the Talbot-Lau interferometer arrangement has for some time now been receiving a great deal of attention. Thus, for example, it is described in the publication by F. Pfeiffer et al. [1], “Hard X-ray dark-field imaging using a grating interferometer”, Nature Materials 7, pages 134 to 137, that the use of X-ray optical gratings permits on the one hand the acquisition of phase-contrast X-ray images that provide additional information about an examination object. On the other hand it is also possible to use not just the phase information but also the amplitude information of scattered radiation for imaging. This enables an imaging modality to be created which is based exclusively on the scatter fractions of the X-ray radiation diffracted by the examination object, in other words a small-angle scattering technique. As a result, very small differences in density in the examination object can be visualized at very high resolutions. Similar findings are also revealed in Joseph J. Zambelli et al. [2], “Radiation dose efficiency comparison between differential phase contrast CT and conventional absorption CT”, Med. Phys. 37 (2010), pages 2473 to 2479.
The wave nature of particles such as X-ray quanta allows phenomena such as diffraction and reflection to be described with the aid of the complex diffraction indexn=1−δ+iβ, where the imaginary part β describes the absorption on which contemporary clinical X-ray imaging such as e.g. that of computed tomography, angiography, radiography, fluoroscopy or mammography is based, and the real part δ describes the phase shift that is considered in differential phase imaging.
An X-ray image acquisition system is known from DE 10 2010 018 715 A1, wherein in order to achieve high-quality X-ray imaging use is made of an X-ray image acquisition system for phase contrast imaging of an examination object which has at least an X-ray emitter having a plurality of field-emission X-ray sources for emitting coherent X-ray radiation, an X-ray image detector, a diffraction grating G1 arranged between the examination object and the X-ray image detector, and a further grating G2 which is arranged between the diffraction grating G1 and the X-ray image detector.
An X-ray image acquisition system by which differential phase contrast imaging of the type cited in the introduction can be performed is known for example from U.S. Pat. No. 7,500,784 B2, which system is explained with reference to FIG. 1.
FIG. 1 shows the typical essential features of an X-ray image acquisition system for an interventional suite, comprising a C-arm 2 which is supported by a stand 1 in the form of a six-axis industrial or articulated-arm robot and at the ends of which an X-ray radiation source, for example an X-ray emitter 3 comprising X-ray tube and collimator, and an X-ray image detector 4 are mounted as the image acquisition unit.
By way of the articulated-arm robot, known for example from U.S. Pat. No. 7,500,784 B2, which preferably has six axes of rotation and consequently six degrees of freedom, the C-arm 2 can be moved arbitrarily in three dimensions, for example by its being rotated about a center of rotation between the X-ray emitter 3 and the X-ray detector 4. The inventive angiographic X-ray system 1 to 4 can be rotated in particular about centers of rotation and axes of rotation in the C-arm plane of the X-ray image detector 4, preferably about the axes of rotation intersecting the center point of the X-ray image detector 4 and the center point of the X-ray image detector 4.
The known articulated-arm robot has a base frame which is permanently installed on a floor, for example. Secured thereto is a turntable which is rotatable about a first axis of rotation. Attached to the turntable so as to be capable of pivoting about a second axis of rotation is a robotic floating link to which a robotic arm is fixed so as to be rotatable about a third axis of rotation. A robotic hand is attached to the end of the robotic arm so as to be rotatable about a fourth axis of rotation. The robotic hand has a securing element for the C-arm 2 which can be pivoted about a fifth axis of rotation and rotated about a sixth axis of rotation running perpendicularly thereto.
The implementation of the X-ray diagnostic apparatus is not dependent on the industrial robot. Conventional C-arm devices can also be used.
The X-ray image detector 4 can be a flat semiconductor detector, rectangular or square in shape, which is preferably produced from amorphous silicon (a-Si). However, integrating and possibly counting CMOS detectors can also be used.
A patient 6 to be examined is positioned as the examination object in the beam path of the X-ray emitter 3 on a tabletop platform 5 of a patient examination table. Connected to the X-ray diagnostic apparatus is a system control unit 7 having an image system 8 which receives and processes the image signals of the X-ray image detector 4 (operator control elements, for example, are not shown). The X-ray images can then be studied on displays of a monitor array 9. Also provided in the system control unit 7 is a known device 10, the function of which will be described in greater detail.
Instead of the X-ray system having the stand 1 in the form of the six-axis industrial or articulated-arm robot shown by way of example in FIG. 1, the angiographic X-ray system can also have, as illustrated in simplified schematic form in FIG. 2, a normal ceiling- or floor-mounted support for the C-arm 2.
Instead of the C-arm 2 shown by way of example, the angiographic X-ray system can also have separate ceiling- and/or floor-mounted supports for the X-ray emitter 3 and the X-ray image detector 4 which are coupled for example in an electronically rigid manner.
In the arrangements for clinical phase contrast imaging that are the focus of attention today, use is made of conventional X-ray tubes, currently available X-ray image detectors, as described for example by Martin Spahn [3] in “Digitale Röntgenbilddetektoren in der Röntgendiagnostik” (“Digital X-ray image detectors in X-ray diagnostics”), Radiologe 43 (2003), pages 340 to 350, and three gratings G0, G1 and G2, as will be explained in more detail hereinbelow with reference to FIG. 2, which shows a schematic layout of a Talbot-Lau interferometer for differential phase contrast imaging with extended tube focus, gratings G0, G1 and G2, and pixelated X-ray image detector.
In order to generate coherent radiation, the X-rays 12 emitted from a tube focus 11 of the non-coherent X-ray emitter 3 penetrate an absorption grating 13 (G0), which effects the spatial coherence of the X-ray radiation source, as well as an examination object 14, the patient 6 for example. In passing through the examination object 14 the wave front of the X-rays 12 is deflected by phase shifting, as illustrated by the normal 15 of the wave front without phase shift, i.e. without object, and the normal 16 of the wave front with phase shift. The phase-shifted wave front then travels through a diffraction or phase grating 17 (G1) having a grating constant adapted to the typical energy of the X-ray spectrum for generating interference lines and in turn through an absorbing analyzer grating 18 (G2) for reading out the generated interference pattern. The grating constant of the analyzer grating 18 is correlated with that of the phase grating 17 and the remaining geometry of the arrangement. The analyzer grating 18 is arranged e.g. at the first or n-th Talbot distance. In this case the analyzer grating 18 converts the interference pattern into an intensity pattern which can be measured by the detector. Typical grating constants for clinical applications are in the region of a few μm, as is also supported for example by the cited literature references [1, 2].
If the tube focus 11 of the radiation source is sufficiently small and the generated radiated power is nonetheless sufficiently great, it may be possible to dispense with the first grating G0, the absorption grating 13, as is the case for example when a plurality of field-emission X-ray sources are provided as the X-ray emitter 3, as is known from the below-described DE 10 2010 018 715 A1.
The differential phase shift is now determined for each pixel of the X-ray image detector 4 in that by way of what is termed “phase stepping” 19, as indicated by an arrow, the analyzer grating 18 G2 is shifted in multiple increments by a corresponding fraction of the grating constant vertically with respect to the radiation direction of the X-rays 12 and laterally with respect to the arrangement of the grating structure and the signal Sk resulting for this configuration during the image acquisition in the pixel of the X-ray image detector 4 is measured and hence the resulting interference pattern sampled. The parameters of a function describing the modulation (e.g. sinusoidal function) are then determined for each pixel by means of a suitable fit method, an adaptation or compensation method, to the thus measured signals Sk. The visibility, i.e. the normalized difference formed from maximum and minimum signal, is in this case a metric for characterizing the quality of a Talbot-Lau interferometer. It is defined as the contrast of the sampled modulation
  V  =                              I                      ma            ⁢                                                  ⁢            x                          -                  I                      m            ⁢                                                  ⁢            i            ⁢                                                  ⁢            n                                                I                      ma            ⁢                                                  ⁢            x                          +                  I                      m            ⁢                                                  ⁢            i            ⁢                                                  ⁢            n                                =                  A                  I          _                    .      
Also in this equation, A denotes the amplitude and Ī the average intensity. The visibility can assume values between zero and one, since all variables are positive and Imax>Imin. In a real interferometer it also holds that Imin>0, thereby beneficially exhausting the value range of V. Minimum intensities greater than zero and all non-ideal characteristics and defects of the interferometer lead to a reduction in visibility. A third item of information which can be defined via the visibility and is generated by this type of measurement is referred to as a dark field. The dark field specifies the ratio formed from the visibilities of the measurement with object and those without object.
  D  =                    V        obj                    V        ref              =                            A          obj                ·                              I            _                    ref                                      A          ref                ·                              I            _                    obj                    
Three different images can then be generated from the comparison of specific derived variables from the fitted functions for each pixel once with and once without object (or patient):
absorption image,
differential phase contrast (DPC) image and
dark-field image.
When reference is made hereinbelow to image, it may be that what is meant is the triumvirate formed from absorption, DPC and dark-field image.
The realization of the method poses many challenges, but in particular has one quite critical disadvantage:
The analyzer grating 18 G2 must be maneuvered into different positions and an X-ray acquisition must then be performed in each position. It is therefore apparent that such a method is highly unsuitable for moving objects (such as non-anesthetized patients or patient organs, e.g. heart, lung), even if the object moves by only slight distances between the different measurements. Such a layout is likewise unsuitable due to the mechanical displacement of the analyzer grating 18, the phase stepping 19, to allow realtime imaging or imaging at higher image frame rates of e.g. 15 frames per second (fps) or even 60 to 100 fps. 3D imaging, in which X-ray emitter 3 with X-ray tube and X-ray image detector 4 is continuously rotated around the patient 6, is also not possible in this way.
It is described in EP 2 633 813 A1 that an interference pattern is produced by means of a relative rotation of the gratings G1 and G2 through an angle (Θ) in the vertical direction with respect to the grating lines. This results for example in M=5 interference fringes within a pixel structure D. In the corresponding direction, i.e. perpendicularly to the grating lines, the readout of the detector is then also driven with a subpixel size of Dy=D/M. The subpixel structure Dy is therefore in the order of magnitude of ⅕ of the pixel size D.
Also described is an analog readout of the detector by means of a light source/laser whose beam width is given precisely by Dy, where Dy has much greater dimensions than the grating structure.